The present invention relates to an apparatus for powering X-ray tubes, and more particularly, to using a stack of medium-voltage high-frequency inverters.
Typical X-ray generators are devices that supply regulated, high-voltage, DC power to X-ray producing vacuum tubes, as well as power to the tube filament. The high-voltage is applied between the anode and the cathode of the tube. In a conventional X-ray tube, X-rays are produced be generating electrons by thermionic emission from a tungsten filament (cathode). The electrons are then accelerated to an anode (which may be rotated for wear-averaging purposes) to generate the X-rays. The X-ray emissions are controlled by the applied voltage or potential between the anode and the cathode, as well as by the anode current. X-ray energy is controlled by the applied voltage, typically between 40 kV and 150 kV for medical applications, but sometimes as low as 20 kV as in mammography.
X-ray intensity is determined by anode current, which is controlled by varying filament power. Varying filament power changes the filament temperature, thus varying the electron emission of the cathode. Most electrons emitted by the cathode reach the anode and constitute anode current. Filament power ranges from a few tens to a few hundred watts. Typically anode currents range from as low as 500 microamperes, as in lower power continuous fluoroscopy, to as high as 1 ampere, as in conventional radiography or during cine-radiography runs and computed-tomography (CT) scans.
The high-voltage is either applied continuously, though at low power levels, or as medium to high power pulses or pulse trains. In continuous mode, typical power levels are on the order of 1 kilowatt (100 kVxc3x971 ma); in pulsed mode, instantaneous power levels are on the order of 150 kW (150 kVxc3x971A). X-ray generators used in medial applications have power ratings in the 10 to 100 kilowatt range.
X-ray generators pypically employ one of two fundamental methods to produce the required high-voltage, DC power. In one method, line frequency generators use a step-up transformer to raise the AC line voltage to the desired level, and then rectify and filter the high AC voltage to obtain DC voltage. Due to the low line frequency and high power levels involved, and due to the high amount of insulation required, the transformer and filter capacitors are very bulky and very expensive. The use of dielectric insulated oil is mandatory to achieve the level of insulation required and to assist in dissipating the heat lost in the transformer windings and other components. The insulating transformer oil creates a large space requirement, creates very heavy equipment, and requires seals which often allow the transformer oil to leak and create an environment hazard as well as degrade the line frequency generator. The second method of X-ray generator involves a high-frequency generator using a high-frequency inverter typically made up of a high-frequency oscillator, a high-frequency high-voltage transformer, a high-frequency high-voltage rectifier, and a high-frequency high-voltage filter to obtain the high DC voltage required. The inverter is powered directly from a low voltage DC source such as a battery bank or from the rectified and filtered AC line. Although many inverter topologies exist, high frequency generators typically use a resonant-inverter topology. In this configuration, the high frequency oscillator drives the primary winding of the transformer through a damped resonant circuit. This resonant circuit is generally composed of an inductor, a capacitor, and an equivalent resistance due to the external load connected to the secondary winding of the transformer, and reflected to the primary. The resonant circuit can be configured with the inductor, capacitor, and resistor in parallel or series. Power transferred to the load, thus voltage across the load, can be varied by changing the oscillator frequency. Power is maximum when the circuit is at resonance, that is, when the inductive reactance is exactly cancelled by the capacitive reactance of the circuit. Power drops when the oscillator frequency is either lower or higher than resonant frequency. High-frequency generators are much smaller and lighter than comparable line frequency generators, due to the reduced size of the transformers, capacitors, and inductors; however, typical high-frequency generators still require use of dielectric insulating oil to insulate and dissipate heat in the transformer windings and other components.
All X-ray generators use a high voltage divider to measure accurately the high-voltage outputs. The high voltage divider is made up of a string of equal value multimegaohm resistors, the top of which is connected to the high voltage output, the bottom of it going to a voltage sampling resistor, that in turn is connected to the high voltage return which is grounded. Typical divider ratio is 1V: 10 kV and divider current is on the order of 1 milliampere (ma). High voltage dividers have to be frequency-compensated by connecting a small capacitor in parallel with each resistor, such as to maintain divider accuracy and pulse shape integrity when the high voltage is pulsed. Capacitor values must be many times larger than the stray capacitances that exist between the divider sections and the surroundings. High voltage capacitors are costly and large, so a typical capacitor-compensated high voltage divider is a bulky and expensive device.
Instead of using compensating capacitors, high voltage dividers can also be guarded by enclosing each resistor in a cylindrical shield section that is maintained at about the same potential as the enclosed resistor, as disclosed in U.S. Pat. Nos. 5,023,769 and 5,391,977. This potential is obtained through a second resistor string that is not used for measurement. This ensures that essentially no current flows through the inevitable stray capacitances since there is very little potential difference between any resistor of the precision divider and its own guard section.
The high voltage output of the X-ray generator is connected to the X-ray tube anode and cathode by means of a pair of high voltage coaxial cables. Cable lengths range from a few feet to about 50 feet. The inner conductor carries the high tension and is thoroughly insulated from the outer coaxial conductor, which is solidly grounded for safety purposes. Because of their coaxial construction, high voltage cables behave as transmission lines; characteristic impedance is normally 50 ohms and capacitance is on the order of 50 picofarads per foot. Tube arcing between anode and cathode, or between either tube electrode and ground, is a rather frequent occurrence. It is equivalent to a momentary short circuit across the tube end of the high voltage cable. Since the high voltage cable acts as a transmission line, the short circuit typically reflects back all of the energy received from the line. The reflected energy adds to the incoming energy and provokes a very large voltage spike at the generator end of the line. The sum of the high voltage output from the generator and the spike will oscillate between twice the normal high voltage output and some negative value, inverting in fact the polarity of the output, until all of the reflected energy has been damped. Due to the large spike, output components of an unprotected X-ray generator will catastrophically and irreversibly fail when the X-ray tube arcs. Nevertheless, many cost-conscience X-ray high-frequency generators are not protected against tube arcing. Tube arc protection is typically implemented with a specially designed lossy inductor, where the inductance of the device slows the rise time of the fault current, and the resistance of the device damps the reflected energy, as disclosed in U.S. Pat. Nos. 5,241,260 and 5,495,165. Slowing the rise time of the fault current allows time for other protective devices, such as fuses and shutdown circuitry, to take over and limit the value of the fault current to tolerable levels. The damping resistance avoids resonance between the high voltage cable and the large filter capacitors of the generator output. The arc protection inductor is large and expensive since it must be carefully designed to withstand the very strong electromagnetic forces and high voltages that develop across it during the onset of the fault.
Precise control of the voltage and phase of the power supply to an X-ray tube is important to ensure proper imaging for diagnostic purposes and to avoid unnecessary exposure of the patient to X-ray radiation which does not produce a useable image. For example, during a conventional radiographic gastrointestinal analysis, the patient ingests a radioopaque liquid containing barium. When the patient ingests the liquid, the doctor turns on the X-ray generating tube at a low level and positions the patient between the X-ray tube and a fluoroscopic screen. The doctor analyses the patient""s gastrointestinal track while the barium flows through it. When the doctor sees a part of the procedure he/she wants to record, she typically replaces the fluoroscopic screen with a photographic plate and increases the X-ray to a level intense enough to expose the plate.
Typical high voltage generators are available for up to 100 kilowatts for medical applications, because of component limitations, rapidly rising costs of components, and because electromagnetic interference emissions become increasingly more expensive to contain. However, many medical procedures require more power and cannot use existing high-frequency generators. Furthermore, load currents below 10 milliamperes are also hard to achieve for high voltage generators. This value is too high for low-power, continuous fluoroscopy which typically runs at 0.5 to 2.0 milliamperes. This minimum current constraint is mostly dictated by the stability criteria of the voltage control loop which requires some amount of damping in the output circuit. In the specific case of the variable frequency resonant inverter, the frequency range required to control output power over such a large range also limits the practical span of power output available from the generator. Furthermore, while power efficiency for high-frequency, high voltage generators at full output can reach 85%, that is, 15% of the input power is lost as heat, efficiency is generally very poor at low power levels such as used in fluoroscopy. Indeed, power losses in high frequency generators are mostly due to switching losses of the active inverted devices. In particular, in the variable frequency, resonant inverter typology, reactive power remains high in the resonant circuit even when the real power delivered to the load is small and, therefore, switch losses remain consequently high.
In the pulsed fluoroscopy operating mode, the X-ray generator output is repetitively switched on and off, typically in synchronism with an X-ray detecting device such as a video camera coupled to a fluoroscopic imaging intensifier. This mode is widely used since it reduces the X-ray dose by turning on the X-ray source only when the detecting device is ready to acquire a new image, and turning it off while the detecting device is busy processing the acquired image. In pulsed fluoroscopy mode, typical pulse repetition rates range from 10 to 90 pulses per second and typical pulse widths from 1 to 10 milliseconds. A 100 microsecond rise time represents 10% of the pulse width of a 1 millisecond pulse. Repetitively pulsing an X-ray generator output on and off means that the output of a generator goes from zero to approximately 100 kV and back to zero, for each pulse. The X-ray tube current will also be pulsed-shaped, its peak value being determined by the generator output voltage and by the tube filament temperature set by filament current. Typical peak currents range from 5 ma to 50 ma.
As mentioned above, the high voltage output of the X-ray generator is connected to the X-ray tube by means of a pair of high voltage coaxial cables that have a capacitance on the order of 50 picofarads per foot and links ranging from 3 to 50 feet. Thus, the total capacitance of these cables must be charged to the full output voltage, and discharged back to zero, for every pulse of the generator output. The charge current must be supplied by the generator and its intensity determines the rate of rise of the voltage by the well-known formula i=C dv/dt and consequently the rise time. In this same manner, the discharge current must flow through a cable discharge circuit that shunts the cable capacitance and is triggered every time discharging is required. The intensity of this discharge current then determines the fall time of the pulse. For example, a 20 foot cable will have a capacitance of around 1,000 picofarads and the charge, or discharge, current will need to be 1,000 milliamps to raise, or drop, the voltage by 100 kV in 100 microseconds.
Pulsed fluoroscopy presents a special challenge to the designers of control loops for all types of X-ray generators in the sense that instantaneous voltage, current, and power vary very widely and very rapidly during each of the repetitive pulses. In the high frequency generator that uses the resonant inverter topology, where output power is solely controlled by varying the oscillator frequency, this is nearly impossible to do neatly without a further discharge module with high voltage and low voltage portions in a circuit to limit the xe2x80x9ctailxe2x80x9d in the output waveform and increase image quality by discharging any capacitive voltage remaining on the cable connected the same, as disclosed in U.S. Pat. No. 5,056,125.
In one illustrative embodiment, a power supply for a device which has a load is disclosed. A first resonant generator and a second resonant generator are connected in parallel and each have a phase output. A control circuit is coupled to the first and second generators and controls the first and second phase outputs. The first and second phase outputs are summed to provide a variable power supply to the load.
In another illustrative embodiment, a method controls first and second generators connected in parallel. The generators each have a phase output. The method includes the steps of setting a first phase to the output of the first generator and setting a second phase to the output of the second generator. The method further includes the step of selectively shifting at least one phase output of the generators to achieve a predetermined magnitude of a voltage in a predetermined time.
In still another embodiment, an apparatus for supplying operating power to an X-ray generating source is disclosed. A frequency oscillator mechanism generates an oscillator frequency. A plurality of magnetically coupled inverter modules are coupled to a plurality of resonant circuits. The resonant circuits include an inductor mechanism, a voltage limiting mechanism, and a resistor mechanism. The apparatus also includes an X-ray generating source.
In yet another illustrative embodiment, an apparatus for supplying operating power to a load device is disclosed. A frequency oscillator mechanism generates an oscillator frequency. A plurality of magnetically coupled inverter modules receive the oscillator frequency and are coupled to a plurality of resonant circuits. The resonant circuits include an inductor mechanism, a voltage limiting mechanism, and a resistor mechanism. The plurality of inverter modules are each coupled to at least one transformer device having a primary winding and a secondary winding. The apparatus also includes at least one DC voltage rectifier mechanism and a load sharing mechanism which substantially equalizes the power load on each inverter module.
In another illustrative embodiment, an apparatus for supplying operating power to a load device is disclosed. A frequency oscillator mechanism generates an oscillator frequency. A plurality of magnetically coupled inverter modules receive the oscillator frequency and are coupled to a plurality of resonant circuits. The resonant circuits include an inductor mechanism, a voltage limiting mechanism, and a resistor mechanism. The plurality of inverter modules are each coupled to at least one transformer device having a primary winding and a secondary winding. The apparatus also includes at least one DC voltage rectifier mechanism and a means for sharing the load of the load device substantially equally between the plurality of resonant inverter modules.
In still another embodiment of a system for generating X-ray beams utilizing a plurality of inverter modules, a method for controlling power is disclosed. The method includes the steps of arranging the plurality of inverter modules interconnected by at least one magnetic coupling and selectively disconnecting at least one module from the magnetic coupling.
In another embodiment of a system for generating X-ray beams utilizing a plurality of inverter modules, a method for dissipating is disclosed. The method includes the steps of arranging the plurality of inverter modules interconnected by at least one magnetic coupling and selectively disconnecting at least one module from the magnetic coupling.
In yet another embodiment, an apparatus supplying operating power to an X-ray generating source is disclosed. A transistor switching circuit includes a slew rate detecting circuit, a variable delay circuit, and a feed back loop coupling the slew rate detecting circuit to the variable delay circuit.
In a further embodiment, a method for switching in a system for generating X-ray beams is disclosed. The method includes the steps of sensing a slew rate and generating a delay time for switching transistors based on the slew rate. The transistors invert a current. The method also includes the step of switching the transistors.
In another embodiment, a shielded resistor divider circuit is disclosed. A resistor mechanism has opposing end terminals and a shield limits electrical noise and stray capacitance from interfering with the operation of the resistor mechanism. The shield includes a plurality of paired conductive members disposed along the length of the resistor mechanism and has opposing end terminals. The pairs of conductive members separate the resistor mechanism into separate portions by providing alternating first and second pairs of conductive members along the length of the resistor mechanism. The shield also comprises a capacitor series comprising a plurality of serially connected capacitor mechanisms disposed a predetermined distance from the resistor mechanism and having opposing end terminals. Each capacitor mechanism is connected between adjacent first and second pairs of conductive members which are connected to the end terminals of the capacitor series. The dynamic impedance of the capacitor series is less than the dynamic impedance of the resistor mechanism. The shield also includes a diode bridge series which has a plurality of connected diode bridges coupled to the capacitor series. The end terminals of the resistor mechanism are connected between a higher-voltage potential and a lower voltage potential. The end terminals of the capacitor series are connected between the higher voltage potential and ground. Electrical noise and stray capacitance is coupled to the capacitor series and does not interfere with the resistor mechanism.
In still another embodiment, a shielded resistor circuit is disclosed. A first insulating sheet and a second insulating sheet each have an opposing inner and outer face. The sheets are disposed in parallel with their inner faces adjacent one another. A resistor mechanism is disposed between the inner faces. A first series of paired conductive members are disposed adjacent the resistor mechanism. Each pair of the first series of conductive members include a first member disposed on one of the inner faces, a second member disposed on the outer face opposing the other inner face, and means for connecting the first and second members. A second series of paired conductive members are disposed adjacent the resistor mechanism. Each pair of the second series of conductive members include a third member disposed on the outer face opposing the one inner face, a fourth member disposed on the other inner face, and means for connecting the third and fourth members. Each pair of the second series are disposed between two pairs of the first series. The combined first and second series have opposing end terminals. A capacitor series, including a plurality of serially connected capacitor mechanisms, is disposed between the inner faces a predetermined distance from the resistor mechanism and has opposing end terminals. Each capacitor mechanism is connected to an adjacent pair of the first series and a pair of the second series. The end terminals of the conductive members are connected to the end terminals of the capacitor series. The dynamic impedance of the capacitor series is less than the dynamic impedance of the resistor mechanism. The shielded circuit also comprises a diode bridge series comprising a plurality of connected diode bridges coupled to the capacitor series. The end terminals of the resistor mechanism are connected between a higher-voltage potential and a lower-voltage potential. The end terminals of the capacitor series are connected between the higher-voltage potential and ground. Electrical noise is coupled to the capacitor series and does not interfere with the resistor mechanism.
In yet another embodiment, an X-ray generating source and regulated power supply is disclosed including an X-ray generating source and a regulator circuit. The regulator circuit receives an input signal and regulates at least one of a duration and an amplitude of the input signal to produce a high-voltage output signal for operating the X-ray generating source. A protection circuit is disposed between the regulator circuit and the X-ray generating source for limiting a rate of change of a transient voltage spike produced at the source to a predetermined value and protecting the regulator circuit. A plurality of series connected resistor mechanisms are coupled to the source of the transient high-voltage spike. The resistor mechanism has a stray inductance.
In a further embodiment, a transformer device is disclosed. A single core includes a substantially rectangular-shaped ferrite core having four sides. Each side is a section of the rectangular-shaped magnet. Two primary windings are mounted on the core. Each primary winding is on an opposing side of the rectangular ferrite core. Two secondary windings are mounted on the core. Each secondary winding is on one of the same opposing sides of the rectangular-shaped ferrite core.
In another embodiment, a transformer device for an X-ray generating device is disclosed. A single core includes a substantially rectangular-shaped magnet having four sides. Each side is a section of the rectangular-shaped ferrite core. Two primary windings are mounted on the core. Each primary winding is on an opposing side of the rectangular magnet. Two secondary windings are mounted on the core. Each secondary winding is on one of the same opposing sides of the rectangular-shaped ferrite core.
Various embodiments of the present invention provide certain advantages and overcome certain drawbacks of prior devices, systems, and methods. Embodiments of the invention may not share the same advantages and those that do may not share them under all circumstances. This being said, the present invention provides numerous advantages including the advantages of achieving high voltage output in an X-ray generator; use of inexpensive and common parts to decrease the costs of maintenance and manufacture; smaller, lighter, and environmentally reliable generators; generating infinite run time in X-ray generators without over-heating; decrease need for environmental cooling fans; limiting high thermal and electrostatic stress concentrations in a power generator; limiting losses at less than full power in a power generator; increasing the duty factor of a high voltage generator; permitting low current operations including low-power continuous fluoroscopy with a high voltage generator; limiting effects of single-shot and recurrent tube arcing; permitting continuous operation with defective inverter modules; permitting single phase and three-phase AC supply or direct supply from a DC source for a high voltage generator; and generating shorter rise time pulses required by high performance pulsed fluoroscopy.
Further features and advantages of the present invention as well as the structure and method of various embodiments of the present invention are described herein in detail below with reference to the accompanying drawings.